Electrolyte Gated Organic Field-Effect Transistor Sensors Based on Supported Biotinylated...

22/11/12 12.08Electrolyte Gated Organic Field-Effect Transistor Sensors Based on Supported Biotinylated Phospholipid Bilayer Operating in Buffer Solution

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adma201203587 10.1002/adma.201203587

CommunicationsCommunicationsSensorsSensors

M. Magliulo, A.Mallardi, Mohd. Y.Mulla, S. Cotrone, B.R. Pistillo, P. F.Vikholm-Lundin, G.Palazzo, L. Torsi............x‒xx

Anchored,Anchored,biotinylated biotinylated [PE-[PE-Q1:AU]Q1:AU] Definit ion?Definit ion? PLsPLsforming the capturinglayers in an electrolyte-gated organic field-effecttransistor (EGOFET)allow label-freeelectronic specificdetection at aconcentration level of 10nM in a high ionicstrength solution. Thesensing mechanism isbased on a [PE-[PE-Q2:AU]Q2:AU] would "strong"would "strong"or maybe "clear" beor maybe "clear" bebetter here?better here? neatcapacitive effect acrossthe PL layers involvingthe charges of the targetmolecules.

Electrolyte-GatedElectrolyte-GatedOrganic Field-Organic Field-Effect TransistorEffect TransistorSensors BasedSensors Basedon Supportedon SupportedBiotinylatedBiotinylatedPhospholipidPhospholipidBilayerBilayer

Electrolyte-Gated Organic Field-EffectElectrolyte-Gated Organic Field-EffectTransistor Sensors Based on SupportedTransistor Sensors Based on SupportedBiotinylated Phospholipid BilayerBiotinylated Phospholipid Bilayer [PE-Q3:AU][PE-Q3:AU] Tit leTit le

reduced in length sl ight ly to make i t more concise. Ok?reduced in length sl ight ly to make i t more concise. Ok?

By Maria Magliulo1 Antonia Mallardi2 Mohammad Yusuf Mulla1 SerafinaCotrone1 Bianca Rita Pistillo1 Pietro Favia1 Inger Vikholm-Lundin3 Gerardo

Palazzo1,* and Luisa Torsi1,*

1M. Magliulo, M. Y. Mulla, S. Cotrone, B. R. Pistillo, P. Favia, G. Palazzo, L. Torsi,Dipartimento di Chimica, Università degli Studi di Bari Aldo Moro, Via Orabona 4,70126, Bari, Italy; E-mail: [email protected]; [email protected] [PE-[PE-

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Q4:AU]Q4:AU] Please provide academic t i t les for al l authors, i .e., Prof. or Dr.Please provide academic t i t les for al l authors, i .e., Prof. or Dr.

2A. Mallardi, CNR-IPCF, Istituto per i Processi Chimico-Fisici Via Orabona 4,70126, Bari, Italy [PE-Q5:AU][PE-Q5:AU] Please provide academic t i t les for al l authors, i .e., Prof. or Dr.Please provide academic t i t les for al l authors, i .e., Prof. or Dr.

3I. Vikholm-Lundin, VTT Technical Research Centre of Finland, Sinitaival 6, FI-33101 Tampere, Finland [PE-Q6:AU][PE-Q6:AU] Please provide academic t i t les for al l authors, i .e.,Please provide academic t i t les for al l authors, i .e.,

Prof. or Dr.Prof. or Dr.

*Correspondence to: G. Palazzo (E-mail: [email protected]); L. Torsi (E-mail: [email protected])

Received: 08-28-2012 ; Revised: 11-05-2012 ; Published online: MM-DD-YYYY

KeywordsKeywordselectronic sensing, organic field-effect transistors, electrolyte-gated organic field-effect transistors

There is an increasing interest in low cost, timesaving, yet reliable, point-of-care assays. Direct

electronic, label-free transduction of bio-recognition events represents a compelling alternative

offering miniaturization, easy data handling and processing. Low costs and versatility [PE-[PE-

Q7:AU]Q7:AU] Reworded for language clari ty. Original meaning retained?Reworded for language clari ty. Original meaning retained? can be provided if organic electronic

devices such as organic field-effect transistors (OFETs) are used as transducers.[1,2] At the

beginning, OFET sensors mostly involved the detection of volatile chemical analytes,[3,4] while

organic electronics allowed fabrication of sensing circuits on flexible substrates.[5,6] However,

bare OFET-sensor responses are based on weak interactions that are non-specific in nature.

Specificity can be achieved by endowing the OFET with receptor molecules capable of selectively

interacting with given analytes. One of the first successful examples, involving a bi-layer structure

formed by a chiral-active layer deposited on top of an organic semiconductor (OSC), has allowed

electronic, chiral differential detection at an unprecedented low concentration.[7]

Mother Nature is an endless source of specific interactions as recognition events are the basis

of life. Accordingly, the challenge is now to develop OFET bio-sensors by integrating

biomolecules into the transistor device.[8] To date, the three architectures shown on the left side

of Figure 1Figure 1, have been proposed.

Figure 1Figure 1 Schemes of the architectures of OFET biosensors proposed to date: a) bi-layer OFET; b)

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functional biological interlayer (FBI) OFET; c) electrolyte-gated OFET (EGOFET). d) Scheme ofphospholipid BIOEGOFET fabrication. e) Typical source-drain current-voltage characteristics (black curves,left ordinate) and gate currents (red curves, right ordinate) of a BIOEGOFET in presence of [PE-[PE-

Q8:AU]Q8:AU] Definit ion ok?Definit ion ok? phosphate-buffered saline (PBS). [PE-Q9:AU][PE-Q9:AU] Author please ignore this query.Author please ignore this query.

Typesetters: Please ensure this f igure spreads across 1.3 columns.Typesetters: Please ensure this f igure spreads across 1.3 columns.

The first one exploits the already introduced bi-layer structure (Figure 1a) with the receptors

anchored at the OSC surface where an aqueous sample is placed. The use of an ultra-thin

polymer dielectric gate allows for convenient sub-volt operation.[9-11]

A second architecture (Figure 1b) is the functional bio-interlayer (FBI) OFET[12] example,

where the bio-receptor layer is located between the OSC and the gate dielectric, opportunely in

direct contact with the transistor electronic channel. Such an arrangement is the key for

unprecedented sensitivity along with selectivity. A drawback is that the analyte has to percolate

through the OSC to reach the receptors, and while this takes place efficiently for species with low

molecular weights such as insulin, it could be hindered for larger protein complexes or viruses.







A third configuration (Figure 1c) applies to the electrochemical OFET (ECT)[13-15] and the

electrolyte-gated OFET (EGOFET),[16-18] the former being limited to the detection of

electrochemically active species. In EGOFETs, the dielectric gating is achieved through the

formation of a Debye‒Helmholtz double layer (a few angstroms thickness) at the interface

between the electrolyte solution and the OSC layer (capacitance of the order of 10 μF cm－2)

allowing for sub-volt operation.[19] On the other hand, operation in pure water has been generally

proposed to minimize the electrochemical doping of the OSC in EGOFETs.[20] In ECTs and

EGOFETs the bio-receptor layer overlies the transistor channel, thus the architectures in Figure

1a and Figure 1c share the problem of efficiently grafting recognition bio-elements, hydrophilic in

nature, to the OSC [PE-Q10:AU][PE-Q10:AU] correct now?correct now? which is notably highly hydrophobic. Such a process

generally involves the presence of polar anchoring groups such as hydroxyl, carboxyl, and amino

groups, covalently bound to the OSC backbone. Indeed, the chemical synthesis of a poly(3-

hexylthiophene) (P3HT) derivative bearing carboxylic acid moieties was recently proposed to

achieve DNA covalent linkage to an EGOFET configuration.[21] Changes upon hybridization with

complementary strands in the 0.1 μM concentration range were detected in pure water while no

signal could be detected in saline solution, this occurrence clearly limiting the use of EGOFETs

for biosensing in real samples. Besides, the presence of polar groups in the bulk of a OSC, has

an impact on the transistor performance even when blends are used.[22]

This work presents a general platform for EGOFET biosensing demonstrating that proper bio-

functionalization can lead to stable and reliable label-free electronic sensing in electrolyte solution

with physiologically relevant ionic strength. This is achieved by anchoring, just at the OSC

surface, a capturing bio-layer capable of controlling ionic diffusion into the OSC. The EGOFET

biological functionalization is shown in Figure 1d (see Experimental for more details). The OSC

surface modification is achieved by plasma enhanced vapor chemical deposition (PE-CVD)[23,24]

recently proposed for configurations 1a[10] and 1c.[25] In the latter the PE-CVD, allowing

deposition of a few nanometer-thick-coating rich in COOH groups (Step 1, Figure 1d), has

already been proven to minimally affect the EGOFET electronic performances. These COOH

groups are eventually used to stably anchor phospholipid (PL) molecules, to the OSC surface.

PLs have been chosen as they are versatile bio-systems, being the main components of


















biological membranes; they are amphipathic molecules, i.e., endowed with both hydrophilic and

hydrophobic moieties, that spontaneously self-assemble in water, forming bilayers facing the

outer aqueous domains with their hydrophilic moiety (the PL head) and, secluding in the interior of

the bilayer, the two hydrophobic acyl chains (the PL tails). Being formed by an non-polar oil-like

inner part, the PL bilayers hold a capacitance of around 1 μF cm－2.[26] Under suitable conditions,

the bilayers in water close themselves forming spherical PL shells known as vesicles. Such

vesicles, containing a fraction of PL-bearing amino groups or biotin (biotinylated PL), are prepared

and immobilized (step 2) at the OSC surface through the formation of amide bonds between the

COOH and the PLs amino groups. Eventually, the anchored vesicles spontaneously fuse,

leading to the formation of a PL bilayer stably grafted at the OSC surface (step 3 of Figure 1d).

The non-polar nature of the deposited PL bilayer interior minimizes ionic diffusion through the

membrane, eventually limiting the OSC doping,[27] while the biotinylated PLs conveniently furnish

the binding sites for streptavidin- or avidin-capturing proteins.

Such a bio-modified EGOFET (hereafter called BIOEGOFET) retains good field-effect

performance, as shown by the current‒voltage characteristics (Figure 1e) measured in

phosphate- buffered saline (PBS) solution (10 mM, pH = 7.4) with high ionic strength (0.15 M).

The source-drain current IDS exhibits well shaped linear and saturation regions as the source-

drain bias (VDS) is swept, showing also good modulation with the gate bias VG. Field-effect

mobility, μFET, is in the 10－3 cm2V－1s－1 range (7.2 × 10－4 ± 4 × 10－4, averaged over 15

samples), while the on/off ratio and the threshold voltage are 204 ±91 and －0.03 ± 0.02 V

respectively. These data demonstrate how the P3HT EGOFET electronic performances have not

been significantly degraded by anchoring the PL bilayer. More importantly, gate leakage currents

are low being in the nA range that, along with the very low hysteresis exhibited by the double run

I‒V curves measured in PBS, show that the PL layer efficiently acts as barrier towards

semiconductor doping overcoming the already identified issue of EGOFET operating in PBS.[21]

Such a bio-electronic platform was subsequently used as sensing device. The binding of the

streptavidin (SA) protein to biotin was chosen as a testbed. Biotin is a small molecule with an

astonishingly high affinity (fM dissociation constant)[28] for the four binding sites of streptavidin or

equivalently, of avidin, both being tetrameric proteins. Accordingly, the electronic performances of

BIOEGOFETs prepared with biotinylated PL bilayers exposed to streptavidin solutions have been









investigated. After the incubation with SA the device was first rinsed to remove unbound proteins

and finally measurements were performed in PBS electrolyte solutions. Representative results are

shown in Figure 2Figure 2a where the BIOEGOFET is exposed to 100 μg/mL (1.6 μM) of streptavidin in

PBS (10 mM, pH = 7.4). A sizable increase of IDS, when compared to that measured in buffer

alone, is seen with the fractional increase at VG =‒0.5 V reaching 60%. In this case the electrolyte

Debye length shortening in saline solution does not appear to impede transduction. Notably, a

comparable response is measured also in a 10-fold diluted PBS solution. While contributions can

be found dealing with OFET biosensing operation in an electrolyte solution,[18,29] to the best of our

knowledge, this is the first report of electronic (not electrochemical) biosensing performed in a

solution containing a physiologically relevant concentration of an electrolyte.

Figure 2Figure 2 a) Transfer characteristics of a BIOEGOFET with biotinylated PLs exposed to PBS 10 mM, pH7.4 (open symbols) and streptavidin (full symbols) solutions. A diagram for the rationale leading to the IDScurrent increase is provided in panels (b) and (c). In panels (d) and (e) negative control experiments areproposed. d) Transfer characteristics of a biotin-free BIOEGOFET exposed to PBS (open circles) and tostreptavidin 100 μg/mL (full triangles) are shown. e) Biotin-bearing BIOEGOFET exposed to bovine albuminserum 1 mg/mL (full diamonds) and to PBS (open circles) is shown.






The sensing mechanisms leading to the observed current increase can be explained by

considering the binding of a negatively charged species (as streptavidin is at pH = 7.4) to a

dielectric interfaced to a p-type OSC. In such an EGOFET, a negative gate bias allows the

accumulation of anions at the electrolyte‒PL interface, while a capacitive coupling induces

positive charge accumulation at the OSC surface (Figure 2b). When the negatively charged SA

proteins bind to the biotin sites at the PL bilayer, they act as an extra (negative) gating inducing

more holes in the transistor channel (Figure 2c); the estimated charge that is added per unit

surface upon formation of the SA‒biotin complex, is －3.578 10－2 C cm－2.[30] Eventually, a

current increase is measured for a [PE-Q11:AU][PE-Q11:AU] Correct?Correct? clear capacitive effect, previously reported

also for the SA binding to a biotinylated p-type silicon nanowire FET.[31] As a further support to

the model, a charge density increase of ca. + [PE-Q12:AU][PE-Q12:AU] correct?correct? 9 × 10－2 C cm－2, [PE-[PE-

Q13:AU]Q13:AU] correct now?correct now? comparable to the estimated added negative charge, is seen already at VG =

0, namely just in the presence of the bound SA. Since the isoelectric point for SA immobilized on

PL monolayer is between 5‒5.5,[32] this detection scheme is expected to work at pH values larger

than 6.

Several control experiments were carried out demonstrating that the transistor response is

specific for SA-biotin binding. BIOEGOFETs bearing no biotin, do not, in fact, respond to the

loading with SA (see Figure 2d). This rules out any contribution of non-specific binding, and

demonstrates that the electronic response is actually due to the anchoring of SA to the

biotinylated PL. To evaluate the selectivity, bovine serum albumin (BSA) was assayed in the

presence of biotinylated PLs. BSA is a protein with a strong tendency to non-specifically adsorb to

surfaces. In Figure 2e the transfer characteristics of the BIOEGOFET, including the supported

biotinylated bilayer, exposed to a BSA solution (1 mg/mL) are compared to the response of the

lone buffer: BSA does not induce any significant variation. Therefore, the SA-induced increase in

IDS (Figure 2a) is the electronic transduction of the highly specific biotin‒streptavidin binding

event.

As a support to the BIOEGOFET unconventional approach, the selective protein binding to the

biotin moiety in the PL layer was further investigated by means of surface plasmon resonance

(SPR). SPR chips were prepared as the BIOEGOFET example and avidin binding was monitored

along with BSA for control purposes. The optimization of the capturing layer density was achieved











by studying the coverage as a function of the surface density of biotin moieties (Figure 3Figure 3a). For

one biotinylated PL every 100 PLs the experimentally bound avidin surface coverage is 150 ± 20

ng/cm2 (2.2 ± 0.3 [PE-Q14:AU][PE-Q14:AU] correct now? correct now? pmol/cm2); this is in good agreement with the

theoretical coverage (1 [PE-Q15:AU][PE-Q15:AU] correct now? correct now? pmol/cm2) estimated assuming a PL head area of

0.7 nm2.[33] Further loading with biotinylated PLs, increases the avidin surface coverage that

eventually saturates at 270 ± 30 ng/cm2 (4.0 ± 0.4 [PE-Q16:AU][PE-Q16:AU] correct now?correct now? pmol/cm2). This is

close to the reported avidin full monolayer coverage on silica (8 [PE-Q17:AU][PE-Q17:AU] correct now?correct now?

pmol/cm2).[34] Accordingly, we have chosen a ratio of one biotinylated PL every 33 PLs for the

electronic sensing. A calibration curve showing avidin surface coverage (as probed by SPR) vs.

the avidin concentration is shown in Figure 3b. A limit of detection (LOD) of a few [PE-[PE-

Q18:AU]Q18:AU] correct?correct? nanomolar is associated with these measurements. In the same figure the

maximum BSA coverage of 20 ± 20 ng/cm2, is shown as a dashed line, indicating that the non-

specific binding is very low in the presence of the anchored PLs bilayer.

Figure 3Figure 3 a) [PE-Q19:AU][PE-Q19:AU] Please note that this f igure wil l be re-inserted by our typesetters in the nextPlease note that this f igure wil l be re-inserted by our typesetters in the next

step of processing so that none of the f igure is corrupted.step of processing so that none of the f igure is corrupted. The avidin surface coverage probed by SPR as afunction of the ratio between biotinylated PL and total PL. b) The avidin surface coverage probed by SPR asa function of the avidin concentration. The PL bilayer has a ratio of biotinylated PL/total PL of 0.03. c)BIOEGOFET response (fractional IDS increment) to streptavidin. Each data point is reported as the averagevalue along with standard error bar estimated over 15 replicates; the line is a logarithmic fit. The inset showsthe same data on a logarithmic plot. [PE-Q20:AU][PE-Q20:AU] Author please ignore this query. Typesetters: Please re-Author please ignore this query. Typesetters: Please re-

insert this f igure ensure that al l axes labels, numbers, and panel labels are not corrupted.insert this f igure ensure that al l axes labels, numbers, and panel labels are not corrupted.







The electronic sensing analytical performances were also investigated. In Figure 3c the

responses (ΔI/I being the fractional IDS increases at VG = －0.5 V) of the BIOEGOFET to different

SA concentrations are shown. For each concentration, measurements were performed on 15

different devices and the average value and standard error (SE) are reported, the latter being

below 20%. The responses clearly scale with the logarithm of SA concentration over two orders of

magnitude and the limit of detection (LOD) is of 10 nM. The LOD was estimated as three times

the standard deviation (over 10 replicates) of the response measured in the PBS reference

solution (see the inset of Figure 3c). This allows label-free detection of biotin in the 10 nM‒1μM

range with a LOD of 10 nM. The responses are also fast as the IDS measurements are acquired

in 45 seconds. The comparison between SPR, a well assessed analytical tool, and the

BIOEGOFET platform shows that similar LODs are obtained. On the other hand, the BIOEGOFET





sensors show performances two orders of magnitude better than those reported for the SA

electronic detection with carbon nanotube FETs (LOD = 2.5 μM).[35] BIOEGOFET sensors,

however, perform less well than the extremely sensitive biotinylated single silicon nanowires FET

whose lowest detection is of tens of [PE-Q21:AU][PE-Q21:AU] correct?correct? picomolar concentration).[31] However,

none of the previous determinations have been proven to deliver a reliable quantification, in a

large concentration range and, critically important, in a high ionic strength electrolyte solution.

Indeed, the BIOEGOFET platform, also matches the performances of customary [PE-Q22:AU][PE-Q22:AU] "label" label

needing" ? Please re-word this sentence.needing" ? Please re-word this sentence. label needing fluorimetric and spectrophotometric assays. In

particular, the LOD lies at the lower edge of the range of assays based on fluorescent biotin

analogues (LOD 30 nM),[36,37] that allow, however, for a much narrower dynamic range.

From an applications point of view, the proposed approach opens a wider perspective as

antibodies and proteins can be also bound to fluid phospholipid bilayers supported on solid

surfaces without altering the membrane properties.[38] The biotinylated PL bilayer anchored to the

OSC surface, can easily be further functionalized with virtually any receptors by using the well-

assessed biotin‒(strept)avidin chemistry.[28] In this respect, the SA-mediated binding of antibodies

to biotinylated-PL-supported bilayer was reported to provide efficient suppression of nonspecific

binding in an acoustic biosensor.[39] Furthermore, while the EGOFET sensors proposed to date

have had reasonable performances only in deionized water, the present BIOEGOFET works

efficiently in high ionic strength electrolyte solution, opening the possibility of detection in real

samples.

ExperimentalExperimental

Materials: Soy bean lecithin (EPIKURON 200) and N-((6-(biotinoyl)amino)hexanoyl)-1,2-

dihexadecanoyl-sn-glycero-3-phosphoethanolamine (Biotin-X DHPE) were respectively

purchased from CARGILL and INVITROGEN. [PE-Q23:AU][PE-Q23:AU] correct now?correct now? Poly-3-hexylthiophene

(P3HT), purchased from BASF (RR>98%, Sepiolid P200), was purified by successive Soxhlet

extractions with methanol and hexane.[40] All other chemicals were purchased from Sigma.

EGOFET fabrication: Si/SiO2 substrates with interdigitated gold source (S) and drain (D)

electrodes were cleaned using solvents of increasing polarity.[41] S and D were












photolithographically defined (10 μm channel length and 10 mm channel width) by using vacuum-

deposited titanium as gold-adhesion-promoter layer. The purified P3HT chloroform solution was

spin-coated on the bottom-contact devices and subsequently annealed at 75 °C for 1h to improve

the OSC performances.[42]

PE-CVD coating: Plasma depositions were carried out in a home-build low pressure RF (13.56

MHz) plasma reactor described elsewhere.[43] Acrylic acid vapors (AA 99%, Aldrich, USA),

ethylene (Eth, Air Liquide) and argon (Ar, Air Liquide) were mixed before being let into the

reactor.[43] Conditions were: power 50 W, pressure 250 mTorr, flow rate of precursors: ΦAA = 2.5

sccm, ΦEth = 7.5 sccm, ΦAr = 2.5 sccm, deposition time was 3 s. PE-CVD conditions were

optimized to the best stability of COOH-functionalized coatings in water media.[44]

Vesicles preparation and PL layer deposition: soybean lecithin (4.5 mg),

phosphatidylethanolamine (0.5 mg), and the proper amount (50‒150 μg) of biotin-X DHPE were

dissolved in chloroform, then dried under N2 flow and kept for 60 min under vacuum. PLs were

rehydrated in PBS (1 mL), sonicated (10 min) on ice, and extruded through a polycarbonate filter

with 100 nm pore sizes using the Avanti mini-extruder (Avanti Polar).

Bio-conjugation: PL vesicles were immobilized onto PE-CVD-modified EGOFETs through the

formation of [PE-Q24:AU][PE-Q24:AU] correct now?correct now? amide bonds between the COOH surface groups and the

NH2 moieties on the polar heads of phosphatidylethanolamine. The EGOFET or SPR gold slides

coated with plasma-treated P3HT were incubated in a freshly prepared

ethyl(dimethylaminopropyl) carbodiimide (EDC) and N-Hydroxysuccinimide (NHS) solution (both

100 mM in PBS) for 1 h at room temperature, rinsed three times in PBS and immersed in the

vesicles suspension overnight under mild agitation at room temperature.

SPR measurements: SPR measurements were performed with a Biacore 3000 instrument (GE

Healthcare/Biacore AB). Substrates coated with lipids were glued with double-sticking tape in a

plastic frame of a Biacore 3000 chip cassette. BSA and avidin were immobilised by flow injection,

after docking the chip and performing a single prime operation to stabilize the surface in PBS. An

approximation that 10 resonance unit (RU) corresponds to 1 ng/cm2 of adsorbed protein was

used.[45]

BIOEGOFET electrical and sensing measurements: The BIOEGOFET current-voltage








BIOEGOFET electrical and sensing measurements: The BIOEGOFET current-voltage

characteristics were measured using a semiconductor parameter analyzer (Agilent 4155C). A

gold plate (1 mm × 1 mm) in contact with a droplet (2 μL) of water or PBS was used as gate

electrode and the measurements were performed in a water-vapor-saturated environment. The

transfer characteristics (IDS vs VG) were measured keeping the drain voltage (VDS) constant at

－0.5 V. The transistor electrical figures of merit, namely the field-effect mobility (μFET), the

threshold voltage (VT) and on/off ratio, were extracted from the characteristics in the saturation

regime.[46] Sensing measurements were performed by measuring the BIOEGOFET transfer

characteristics in PBS (10 mM, KCl 2.7 mM, 137 mM NaCl, pH = 7.4). Solutions of SA (or BSA) in

buffer were placed on the S-D channel for 15 minutes. The channel was then rinsed with buffer to

remove the unbound proteins and finally the response was evaluated as the IDS relative current

increase at VG = －0.5V.

Supporting InformationSupporting Information

Supporting Information is available from the Wiley Online Library or from the author.

AcknowledgementsAcknowledgementsThe “Electrolyte-Gated Organic Field-Effect Biosensors- BioEGOFET” project, SEVENTHFRAMEWORK PROGRAMME - FP7/2007-2013 under Grant agreement no. 248728, and the“Gas Sensors on Flexible Substrates for Wireless Applications-FlexSMELL project, SEVENTHFRAMEWORK PROGRAMME‒PITN/2007-2013 under Grant agreement no. 2009-238454 areacknowledged for partial financial support. We also acknowledge funding from The Italian Ministryof Education, University and Research (MIUR) through PRIN project “Electronic andelectrochemical biosensors” 2009AZKNJ7. Dr. Francesca Intranuovo is acknowledged forperforming some PE-CVD deposition experiments. Henrick Toss from the University of Linkopingis acknowledged for supplying the photolithographic patterning of the S&D pads in the EGOFETdevices. Prof. Gilles Horowitz of LPICM, Ecole Polytechnique, Palaiseau, France, and Prof.Magnus Berggren from the University of Linkoping are acknowledged for useful discussions.

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